1. Field of the Invention
The present invention is directed to a method for the operation of a magnetic resonance apparatus with a gradient system containing at least one gradient coil for generating a gradient field at least within an imaging volume and at least one shield coil operable independently of the gradient coil for generating a shielding field with which the gradient field can be counteracted in a prescribable region, and is also directed to a magnetic resonance apparatus for the implementation of the method.
2. Description of the Prior Art
Magnetic resonance technology is a known technique for acquiring images of the inside of the body of an examination subject. Rapidly switched gradient fields that are generated by a gradient system are superimposed in a magnetic resonant apparatus on a static basic magnetic field that is generated by a basic field magnet system. The magnetic resonance apparatus further has a radio-frequency system that emits radio-frequency signals into the examination subject for triggering magnetic resonance signals and that registers the generated magnetic resonance signals, from which magnetic resonance images are produced.
The gradient system contains a gradient coil system and a gradient amplifier and control unit. The gradient coil system usually has three gradient coils. Each of the gradient coils generates a gradient field for a specific spatial direction that, in the desired ideal case, is exclusively formed by a primary field component that is collinear with the basic magnetic field at least within an imaging volume. The main field component has a prescribable gradient that, in the desired ideal case, is of the same magnitude independently of location at every arbitrary point in time, at least within the imaging volume. Since the gradient field is a time-variable magnetic field, this in fact applies for every point in time; however, the intensity of the gradient is variable from one point in time to another. The direction of the gradient is usually permanently prescribed by the design of the gradient coil. As a result of Maxwell""s fundamental equations, however, and contrary to the desired ideal case, no gradient coils can be formed that produces only the aforementioned primary field component over the imaging volume. Among other things, at least one accompanying field component that is directed perpendicularly to the primary field component unavoidably accompanies the primary field component.
Appropriate currents are set in the gradient coil for generating the gradient field. The amplitudes of the required currents amount to up to several hundred amperes. The rise and decay rates of the current amount to up to several hundred kA/s. For power supply, the gradient coils are connected to the gradient amplifier and control unit.
The gradient coil system usually is surrounded by conductive structures wherein eddy currents are induced by the activated gradient fields. Examples of such conductor conductive structures are the vacuum vessel and/or cryoshield of the superconducting basic field magnetic system, copper foil of the radio-frequency shielding and the gradient coil system itself. The fields generated by the eddy currents are unwanted because, without counter-measures, they attenuate the gradient field and distort it in terms of its time curve. This leads to degradation of the quality of the magnetic resonance images. Further, the eddy currents induced in the superconducting basic field magnet system cause a heating of the basic magnetic system, so that a considerably increased cooling power must be exerted for maintaining the super-conduction. In the case of a basic field magnetic system with a permanent magnet, the heating as a consequence of the eddy currents leads to an unwanted modification of the properties of the basic magnetic field and, further, the eddy currents can even produce a demagnetization of the permanent magnet.
Such eddy current fields can be compensated to a certain degree by a suitable pre-distortion of a reference current quantity of the gradient coil. With the pre-distortion, however, only eddy current fields can be compensated that image the gradient field similarly in the mathematical sense, i.e. are the same as the gradient field in terms of their spatial course. The principle functioning of such pre-distortion is disclosed in U.S. Pat. No. 4,585,995. The calculation of the pre-distortion is thereby essentially based on the perception that excited and decaying eddy currents can be described by a specific number of exponential functions having different time constants. Transferred to an electrical network for the compensation of eddy current fields, this means that the pre-distortion can be implemented with filters having different limit frequencies. The setting of the time constants or limit frequencies ensues, for example, by an operator who determines the optimum values at the installed magnetic resonance apparatus by step-by-step variation of settings of the pre-distortion and repeated checking. In another embodiment, the setting of the time constants or limit frequencies ensues automatically. The latter is disclosed, for example, in U.S. Pat. No. 4,928,063.
When implementing a sequence, the pre-distortion of the reference current quantity should be continuously implemented during the entire time execution of the sequence. Due to the pre-distortion of the gradient field amplifier and control unit, power reserves must be kept available that generate a higher power and thus resulting in a more costly dimensioning of the gradient amplifier and control unit.
Since, however, the gradient field also produces eddy current fields whose spatial curves are not the same as the gradient field, additional spatial field distortions of a higher order arise. In order to largely compensate these field distortions, actively shielded gradient coils are among the measures utilized. A shield coil belonging to the gradient coil is designed for this purpose such that the gradient field can be neutralized (counteracted) in a prescribable region, usually in a vacuum container surrounding the gradient coil system or a cryoshield of a superconducting basic field magnet system. To this end, the shield coil usually has a lower number of turns than the gradient coil and is interconnected with the gradient coil so that the shield coil has the same current therein as the gradient coil, but flowing in the opposite direction. Further, the shield coil has an attenuating effect on the gradient field in the imaging volume; an attenuation of the actually useful gradient field of up to half in the imaging volume must be accepted. A gradient coil with an appertaining shield coil for neutralizing a gradient field on a defined area is disclosed, for example, in British Specification 2 180 943.
Further, German OS 34 11 222 discloses a magnetic resonance apparatus that has three gradient coils for generating gradient fields and at least one further coil arrangement operable independently of the gradient coils for generating a magnetic field that proceeds in the direction of a basic magnetic field. The further coil arrangement is designed such that the magnetic field changes in a spatially non-linear fashion and such that a superimposition of the magnetic field with the gradient fields yields a defined, time-spatial modification of the magnetic flux density. The further coil arrangement is fashioned such in one embodiment so that the magnetic field has a spatial course that corresponds to a spherical function of the second or third order. In particular, the unwanted eddy current effects caused by the gradient fields can be eliminated with the further coil arrangement.
An object of the present invention is to provide an improved method for operating a magnetic resonance apparatus as well as a magnetic resonance apparatus for the implementation of the method with which, among other things, high gradient intensities can be achieved.
In the inventive method for operating a magnetic resonance apparatus having a gradient system containing at least one gradient coil for generating a gradient field at least within an imaging volume, and at least one shield coil operable independently of the gradient coil for generating a shielding field with which the gradient field can be neutralized in a prescribable region, the above object is achieved by operating the gradient system in an operating mode wherein the shield coil is operated for intensifying the gradient field within the imaging volume.
Extremely high gradient intensities can be achieved in the imaging volume with this operating mode. This is especially advantageous, for example, for producing a diffusion gradient pulse, that should have an extremely high gradient/time integral with an optimally short time duration. The analogous case applies for a spoiler gradient pulse. It is accepted in this operating mode that the gradient field is not neutralized by the shielding field in the prescribable region. If the prescribable region is, for example, a region of a basic field magnet system or a region of a cryoshield of a superconducting basic field magnet system, then an increased heating of the basic field magnet system in this operating mode is accepted in favor of the high gradient intensities. This heating can be compensated, for example, by increasing the cooling power of a cooling device of the basic field magnet system.
In an embodiment, the gradient coil in this operating mode is operatedxe2x80x94with respect to at least one component of eddy current fieldsxe2x80x94free of compensation and the shield coil is simultaneously operated for compensating at least one component of eddy current fields. The relevant component or components of eddy current fields is/are produced by the gradient field and/or the shielding field and occur within the imaging volume. Because the gradient coil is operated free of compensation, i.e. without a pre-distortion, the gradient coil can be operated with its maximally allowed current amplitude for generating high gradient intensities during an overall duration of, for example, a square-wave-shaped gradient pulse. The shield coil is operated such that the gradient field is intensified in the imaging volume and eddy current fields produced by the gradient and shielding field are simultaneously compensated. To that end, the shield coil is operated with a suitable pre-distortion. The pre-distortion is accomplished, for example, by filtering a quantity that controls the shield coil current. For suppressing eddy current influences that disturb magnetic resonance images, it is adequate in many cases for only one component of eddy current fields to be compensated, the spatial course thereof corresponding to a spherical function of the first order.
In another embodiment, the gradient field is operated in aforementioned operating mode for compensating a part of at least one component of eddy currents. Because only one part of a component of eddy current fields, whose spatial curve, for example, corresponds to a spherical function of the first order, is compensated with the gradient coil, the shield coil can be operated such that at least one further component of eddy current fields can be additionally compensated, whose spatial course corresponds to a spherical function of a higher order.
In a further embodiment, the gradient system is operated in a further operating mode wherein the shield coil is operated for neutralizing the gradient field on the prescribable region. To that end, for example, the shield coil is operated with a current of the opposite direction but equal in magnitude with respect to the gradient. When the prescribable region is the aforementioned region of the gradient field magnet system, then a heating of the basic field magnet system is minimized in the further operating mode due to the neutralization. A good compensation of eddy current fields is likewise achieved. However, the gradient field is attenuated in the imaging volume by the shielding field in exchange.
In another embodiment, a switch is made between the operating conditions in a time sequence, for example within the framework of a sequence. The first operating mode is used during a time segment of the sequence wherein high gradient intensities are required, for example, for diffusion gradients, and the second operating mode is otherwise used. Further, the gradient system can be operated in an at least one further operating mode between the aforementioned first and the second operating modes. Such an intermediate mode can require in that the first operating mode be set for a calculated gradient axis and the second operating mode is simultaneously set for a further calculated gradient axis. If the calculated gradient axes are unequal to the physical gradient axes permanently prescribed by the gradient system, then a conversion from calculated to physical gradient axes automatically leads to the operation of the gradient system with the aforementioned intermediate mode.